The invention relates to a method for measuring the cerebral perfusion of a living organism by means of magnetic resonance (=“MR”) imaging, wherein the change in the MR signal is measured during the passage of a contrast medium bolus in a measurement volume in the brain of the living organism and wherein, additionally in a slice that is pervaded by an artery that supplies the blood to the brain, further MR signals are acquired to determine the temporal progression of the concentration of the contrast medium in the artery during the bolus passage.
Such a method is known from “Measuring the Arterial Input Function With Gradient Echo Sequences” by Matthias J. P. van Osch, Evert-jan P. A. Vonken, Max A. Viergever, Jeroen van der Grond, and Chris J. G. Bakker, Magnetic Resonance in Medicine, 49:1067-1076 (2003).
The object of a dynamic susceptibility contrast (=DSC) measurement is to determine perfusion parameters, such as cerebral blood volume (=CBV), cerebral blood flow (=CBF), and mean transit time (=MTT) with local resolution. If a contrast medium bolus is injected into the arm vein, the passage of this bolus through the brain will produce a time-variable contrast in the magnetic resonance (=MR) image. Fast MR sequences (e.g. echo planar imaging (=EPI)), enable temporally and spatially resolved measurement of the passage of this contrast medium bolus. The change in the relaxation rate can be determined for each voxel from the time series measured in this way. The contrast medium concentration can be approximately determined from the change in the relaxation rate.
The above parameters are determined using the tracer kinetic model [1, 2, 3]. This model establishes a connection between the progression of the contrast medium concentration in the arterial supply (cin (t)) and the progression of the contrast medium concentration in the voxel under consideration (ct (t)). The above parameters can be determined by comparing the two progressions.
The signal change caused by the contrast medium is very different in the tissue and in the arteries. The dynamic range of the measurement is so limited that the arteries and tissue cannot be optimally measured with the same echo time (=TE). With a short TE, it is possible to measure the change in relaxation rate in large vessels, while in tissue, where the CBV is small, the effect of the contrast medium is no longer visible. With a long TE, the effect in the tissue is easily visible but the magnetization in large vessels is then almost completely relaxed and therefore does not produce a signal above the noise level. A very short echo time would enable measurement of the arterial blood but is very complex to image and has so far not been implemented. In a standard protocol, a TE is selected that is optimized for tissue. There are approaches in which multiple echoes of the same excitation [4,6] are acquired, but it has been shown that the shortest echo time is still too long to measure the arterial signal.
The fast MR sequences required for perfusion measurements, such as EPI, suffer from an artifact in which the inhomogeneities of the main magnet field result in a disturbance of the spatial encoding due to the associated changes in local Larmor frequency. This is manifested as an artificial shift of image elements [10]. Such inhomogeneities arise locally in the vicinity of blood vessels due to the contrast medium. This is manifested as an apparent movement of arteries through the image during the bolus passage.
The tracer kinetic model requires that Cin be the direct input of the voxel under consideration. As explained above, the arterial input function (=AIF) is determined further away in the vessel tree by way of a substitute. In this additional way, the shape and arrival time of the bolus is changed due to the blood flow conditions. To minimize these effects, it was suggested that an individual AIF be determined for a given brain area instead of a global AIF [5]. Given the relatively low spatial resolution, this is associated with such local AIFs in arteries being determined as very small compared with the voxel size. This so-called partial volume effect results in a severe loss of the arterial contribution in the measured signal.
To minimize partial volume effects, van Osch et al. have proposed in [6] that the complex-value signal be recorded during DSC measurement. If a vessel is chosen that is parallel to the magnetic field, it is possible to separate the contribution of the large vessel from that of the surrounding tissue. However, such a vessel is hard to find. Straight blood vessels that can be used as local AIF are generally not parallel to the magnetic field and are thus not suitable for correction. Large vessels, such as the internal carotid are almost parallel to the magnetic field but the high contrast medium concentration per voxel exceeds the dynamic range of a typical measurement.
The currently most frequently used method of DSC evaluation is based on selection of a global AIF. Selection can be performed manually if the user chooses a voxel whose signal the user considers to be a suitable AIF.
It has been shown that the resulting perfusion parameters heavily depend on the user. To obtain comparable and reproducible results, methods of automatic AIF selection have been suggested [7, 8, 9]. Just like manual selection, an AIF is obtained that is distorted by partial volume effects and apparent movement.
In patent EP 0 958 503 B1, it is suggested that AIF be measured in the neck of the patient. This is done by exciting a one-dimensional volume, e.g. a cylinder that contains the carotids. The selective excitation pulses that this requires are very long, which is why the measurement of the blood signal becomes too inaccurate for high contrast medium concentrations because the TE is too long. Shortening the pulse lengths, although conceivable, would worsen the selection profile, which would result in partial volume problems. Moreover, this solution does not provide a way of handling the artificial shift of the arteries caused by the contrast medium.
Van Osch et al. proposed a further solution and placed an excitation slice through the neck, which was pervaded by the arteries. For each slice, during the brain measurement, van Osch et al. attempted to determine the AIF with a short echo time. A complete 2-D acquisition was made of the neck slice. During the passage of the contrast medium bolus, however, the signal almost completely dropped below the noise level. Moreover, several excitations of the neck slice were required in the brain for each slice because of the short echo time.
The problems with the prior art described above can be summarized as follows:                Very different concentrations of the contrast medium in large vessels and tissue make accurate and simultaneous measurement of large vessels and tissue with the same sequence parameters impossible.        For perfusion measurement, the sequence parameters must be optimized for the tissue contrast. That results in the arterial signal dropping below the noise level.        Arterial input functions that can be determined with the means of the prior art can be distorted by partial volume effects.        The change in the Larmor frequency due to the contrast medium results in a shift in the position of the arteries in the acquired image.        The temporal resolution of the DSC measurement is very low, especially for the AIF.        
The object of this invention is therefore to provide a method that enables simultaneous measurement of the contrast medium flow in vessels supplying the brain and brain tissue within the same sequence with an adapted dynamic range.